Image processing apparatus, radiography system, image processing method, and image processing program

ABSTRACT

An image processing apparatus includes: an acquisition unit that acquires a radiographic image generated by a radiation detector irradiated with radiation from a radiography apparatus including the radiation detector in which plural pixels, each of which includes a conversion element that generates a larger amount of charge as it is irradiated with a larger amount of radiation, are arranged; and a correction unit that corrects scattered ray components caused by scattered rays of the radiation included in the radiographic image, using correction data for correcting scattered rays which is associated with each of plural regions in the radiographic image.

CROSS-REFERENCE TO RELATED APPLICATION

This application claims priority from Japanese Patent Application No.2018-057224, filed on Mar. 23, 2018, the disclosure of which isincorporated by reference herein in its entirety.

BACKGROUND Field of the Invention

The present disclosure relates to an image processing apparatus, aradiography system, an image processing method, and a non-transitorystorage medium storing an image processing program.

Related Art

An image processing apparatus has been disclosed which converts aconverted image obtained by converting a pixel value higher than areference value among a plurality of pixel values forming a medicalimage into a lower pixel value into a scattered ray image in the medicalimage on the basis of a scattering function (see JP2017-164626A). Theimage processing apparatus subtracts the scattered ray image from themedical image to generate a scattered ray reduced image in whichscattered rays have been reduced.

However, in recent years, it is desirable to stably and accuratelycorrect components caused by the scattered rays of radiation included ina radiographic image. For example, in a method for calculating anumerical value from a radiographic image, such as a dual-energy X-rayabsorptiometry (DXA) method for deriving the bone density of a subjectfrom a radiographic image, it is desirable to stably correct componentscaused by the scattered rays of radiation included in a radiographicimage with higher accuracy.

However, in the technique disclosed in JP2017-164626A, a difference inscattered rays between regions is not considered. Therefore, in somecases, it is difficult to correct components caused by the scatteredrays of radiation included in a radiographic image with high accuracy.

SUMMARY

The present disclosure has been made in view of the above-mentionedproblems and an object of the present disclosure is to provide an imageprocessing apparatus, a radiography system, an image processing method,and a non-transitory storage medium storing an image processing programthat can correct components caused by the scattered rays of radiationincluded in a radiographic image with high accuracy.

In order to achieve the object, according to the present disclosure,there is provided an image processing apparatus comprising: anacquisition unit that acquires a radiographic image generated by aradiation detector irradiated with radiation from a radiographyapparatus including the radiation detector in which a plurality ofpixels, each of which includes a conversion element that generates alarger amount of charge as it is irradiated with a larger amount ofradiation, are arranged; and a correction unit that corrects scatteredray components caused by scattered rays of the radiation included in theradiographic image, using correction data for correcting scattered rayswhich is associated with each of a plurality of regions in theradiographic image.

In the image processing apparatus according to the present disclosure,the radiography apparatus may include two radiation detectors, that is,a first radiation detector and a second radiation detector. The firstradiation detector and the second radiation detector may be arrangedalong a direction in which the radiation is emitted. The acquisitionunit acquires a first radiographic image generated by the firstradiation detector irradiated with radiation with a first energy leveland a second radiographic image generated by the second radiationdetector irradiated with radiation with a second energy level differentfrom the first energy level. The correction unit may correct thescattered ray components included in the first radiographic image, usingeach of the plurality of regions and first correction data associatedwith the first radiation detector, and correct the scattered raycomponents included in the second radiographic image, using each of theplurality of regions and second correction data which is associated withthe second radiation detector and is different from the first correctiondata.

In the image processing apparatus according to the present disclosure,the acquisition unit may acquire a first radiographic image generated bythe radiation detector irradiated with radiation with a first energylevel and a second radiographic image generated by the radiationdetector irradiated with radiation with a second energy level differentfrom the first energy level. The correction unit may correct thescattered ray components included in the first radiographic image, usingeach of the plurality of regions and first correction data associatedwith the first energy level, and correct the scattered ray componentsincluded in the second radiographic image, using each of the pluralityof regions and second correction data which is associated with thesecond energy level and is different from the first correction data.

The image processing apparatus according to the present disclosure mayfurther comprise a derivation unit that derives at least one of bonedensity or bone mineral content, using the first radiographic image andthe second radiographic image corrected by the correction unit.

In the image processing apparatus according to the present disclosure,the correction data may include information indicating intensity of thescattered rays and information indicating a spread of the scatteredrays.

In the image processing apparatus according to the present disclosure,the plurality of regions may include a directly irradiated region thatis directly irradiated with the radiation and a subject region that isirradiated with the radiation through the subject in the radiographicimage.

In the image processing apparatus according to the present disclosure,the subject region may include regions corresponding to a plurality ofparts of the subject.

In the image processing apparatus according to the present disclosure,the correction data associated with the subject region may be furtherassociated with a body thickness of the subject.

The image processing apparatus according to the present disclosure mayfurther comprise an estimation unit that estimates the body thickness ofthe subject from a pixel value of the subject region in the radiographicimage.

In the image processing apparatus according to the present disclosure,the correction data may be further associated with imaging conditions.

In the image processing apparatus according to the present disclosure,the imaging conditions may include at least one of a material forming abulb of a radiation source, a tube voltage, a material forming aradiation limitation member, characteristics of a grid, a distance fromthe radiation source to a radiation detection surface of the radiographyapparatus, or a quality of a material forming a radiation incidentsurface of a case accommodating the radiography apparatus.

In the image processing apparatus according to the present disclosure,the correction unit may further correct the scattered ray componentsincluded in the radiographic image, using the correction data associatedwith a distance of a gap between a surface of the subject which facesthe radiography apparatus and the radiation incident surface of the caseaccommodating the radiography apparatus.

The image processing apparatus according the present disclosure mayfurther comprise a derivation unit that derives a first body thicknessof the subject in a portion of the subject which comes into contact withthe radiation incident surface of the case and a second body thicknessof the subject in a portion in which the gap is present, using a pixelvalue of a region corresponding to the subject in the radiographicimage, and derives the distance of the gap, using the first bodythickness and the second body thickness, assuming that the bodythickness of the subject is symmetrically reduced.

In the image processing apparatus according to the present disclosure,each of the first and second radiation detectors may comprise a lightemitting layer that is irradiated with the radiation and emits light.The plurality of pixels of each of the first and second radiationdetectors may receive the light, generate the charge, and accumulate thecharge. The light emitting layer of one of the first and secondradiation detectors which is provided on an incident side of theradiation may include CsI and the light emitting layer of the otherradiation detector may include GOS.

In order to achieve the object, according to the present disclosure,there is provided a radiography system comprising: the image processingapparatus according to the present disclosure; and a radiographyapparatus that outputs a radiographic image to the image processingapparatus.

In order to achieve the object, according to the present disclosure,there is provided an image processing method comprising: acquiring aradiographic image generated by a radiation detector irradiated withradiation from a radiography apparatus including the radiation detectorin which a plurality of pixels, each of which includes a conversionelement that generates a larger amount of charge as it is irradiatedwith a larger amount of radiation, are arranged; and correctingscattered ray components caused by scattered rays of the radiationincluded in the radiographic image, using correction data for correctingscattered rays which is associated with each of a plurality of regionsin the radiographic image.

In order to achieve the object, according to the present disclosure,there is provided a non-transitory storage medium storing a program thatcauses a computer to perform an image processing, the image processingincluding: acquiring a radiographic image generated by a radiationdetector irradiated with radiation from a radiography apparatusincluding the radiation detector in which a plurality of pixels, each ofwhich includes a conversion element that generates a larger amount ofcharge as it is irradiated with a larger amount of radiation, arearranged; and correcting scattered ray components caused by scatteredrays of the radiation included in the radiographic image, usingcorrection data for correcting scattered rays which is associated witheach of a plurality of regions in the radiographic image.

According to the present disclosure, it is possible to correctcomponents caused by scattered rays of radiation included in aradiographic image with high accuracy.

BRIEF DESCRIPTION OF THE DRAWINGS

FIG. 1 is a block diagram illustrating an example of the configurationof a radiography system according to each embodiment.

FIG. 2 is a side cross-sectional view illustrating an example of theconfiguration of a radiography apparatus according to a firstembodiment.

FIG. 3 is a block diagram illustrating an example of the configurationof a main portion of an electric system of a radiography apparatusaccording to each embodiment.

FIG. 4 is a block diagram illustrating an example of the configurationof a main portion of an electric system of a console according to eachembodiment.

FIG. 5 is a graph illustrating the amount of radiation that reaches eachof a first radiation detector and a second radiation detector.

FIG. 6 is a front view illustrating an example of a region from which aDXA profile used to derive bone density is to be derived.

FIG. 7 is a graph illustrating a bone density derivation process.

FIG. 8A is a perspective view illustrating calibration according to eachembodiment.

FIG. 8B is a side view illustrating calibration in a subject regionaccording to each embodiment.

FIG. 8C is a side view illustrating calibration in a directly irradiatedregion according to each embodiment.

FIG. 9A is a graph illustrating an example of correction data accordingto the first embodiment.

FIG. 9B is a graph illustrating an example of the correction dataaccording to the first embodiment.

FIG. 10 is a diagram illustrating a gap between a subject and a caseaccording to each embodiment.

FIG. 11 is a flowchart illustrating an example of an overall imagingprocess according to the first embodiment.

FIG. 12 is a flowchart illustrating an example of an individual imagingprocess according to the first embodiment.

FIG. 13 is a flowchart illustrating an example of a bone densityderivation process according to the first embodiment.

FIG. 14 is a diagram illustrating a process of correcting scattered rayscaused by the direct emission of radiation.

FIG. 15 is a diagram illustrating a process of correcting scattered rayscorresponding to a body thickness of the subject.

FIG. 16 is a graph illustrating an example of a gap deriving profile.

FIG. 17 is a graph illustrating an example of the relationship betweenthe body thickness and a pixel value.

FIG. 18 is a diagram illustrating a process of deriving the distance ofthe gap.

FIG. 19 is a side cross-sectional view illustrating an example of theconfiguration of a radiography apparatus according to a secondembodiment.

FIG. 20 is a graph illustrating the amount of radiation absorbed by aradiation detector in a case in which radiation is emitted at differenttube voltages.

FIG. 21 is a flowchart illustrating an example of an overall imagingprocess according to the second embodiment.

FIG. 22 is a flowchart illustrating an example of an individual imagingprocess according to the second embodiment.

FIG. 23 is a flowchart illustrating an example of a bone densityderivation process according to the second embodiment.

DETAILED DESCRIPTION

Hereinafter, embodiments of the present disclosure will be described indetail with reference to the drawings.

First Embodiment

First, the configuration of a radiography system 10 according to thisembodiment will be described with reference to FIG. 1. As illustrated inFIG. 1, the radiography system 10 includes a radiation emittingapparatus 12, a radiography apparatus 16, and a console 18. The console18 is an example of an image processing apparatus according to thepresent disclosure.

The radiation emitting apparatus 12 according to this embodimentincludes a radiation source 14 that irradiates a subject W, which is anexample of an imaging target, with radiation R such as X-rays. Theradiation emitting apparatus 12 according to this embodiment emits theradiation R with a cone-beam shape. An example of the radiation emittingapparatus 12 is a treatment cart. A method for commanding the radiationemitting apparatus 12 to emit the radiation R is not particularlylimited. For example, in a case in which the radiation emittingapparatus 12 includes an irradiation button, a user, such as a radiologytechnician, may press the irradiation button to command the emission ofthe radiation R such that the radiation R is emitted from the radiationemitting apparatus 12. In addition, for example, the user, such as aradiology technician, may operate the console 18 to command the emissionof the radiation R such that the radiation R is emitted from theradiation emitting apparatus 12.

In a case in which the command to emit the radiation R is received, theradiation emitting apparatus 12 emits the radiation R from the radiationsource 14 according to set emission conditions, such as a tube voltage,a tube current, and an emission period. Hereinafter, the dose of theradiation R is simply referred to as “the amount of radiation”.

Next, the configuration of the radiography apparatus 16 according tothis embodiment will be described with reference to FIG. 2. Asillustrated in FIG. 2, the radiography apparatus 16 includes aplate-shaped housing 21 that transmits the radiation R and has awaterproof, antibacterial, and airtight structure. The housing 21includes a first radiation detector 20A and a second radiation detector20B that detect the radiation R transmitted through the subject W. Inaddition, the housing 21 includes a radiation limitation member 24, acontrol substrate 26A, a control substrate 26B, and a case 28. Theradiography apparatus 16 captures radiographic images of the subject Wusing the first radiation detector 20A and the second radiation detector20B. Hereinafter, in a case in which the first radiation detector 20Aand the second radiation detector 20B do not need to be distinguishedfrom each other, they are generically referred to as “radiationdetectors 20”. In addition, a grid 23 for removing scattered rays isprovided between the housing 21 and the subject W.

The first radiation detector 20A is provided on the incident side of theradiation R and the second radiation detector 20B is provided so as tobe stacked on the side of the first radiation detector 20A from whichthe radiation R is transmitted and emitted. The first radiation detector20A includes a thin film transistor (TFT) substrate 30A and ascintillator 22A which is an example of a light emitting layer that isirradiated with the radiation R and emits light. The TFT substrate 30Aand the scintillator 22A are stacked in the order of the TFT substrate30A and the scintillator 22A from the incident side of the radiation R.The term “stacked” means a state in which the first radiation detector20A and the second radiation detector 20B overlap each other in a casein which the first radiation detector 20A and the second radiationdetector 20B are seen from the incident side or the emission side of theradiation R in the radiography apparatus 16 and it does not matter howthey overlap each other. For example, the first radiation detector 20Aand the second radiation detector 20B, or the first radiation detector20A, the radiation limitation member 24, and the second radiationdetector 20B may overlap while coming into contact with each other ormay overlap with a gap therebetween in the stacking direction.

The second radiation detector 20B includes a TFT substrate 30B and ascintillator 22B which is an example of the light emitting layer. TheTFT substrate 30B and the scintillator 22B are stacked in the order ofthe TFT substrate 30B and the scintillator 22B from the incident side ofthe radiation R.

That is, the first radiation detector 20A and the second radiationdetector 20B are irradiation side sampling (ISS) radiation detectorsthat are irradiated with the radiation R from the side of the TFTsubstrates 30A and 30B.

In the radiography apparatus 16 according to this embodiment, thescintillator 22A of the first radiation detector 20A and thescintillator 22B of the second radiation detector 20B have differentcompositions. Specifically, for example, the scintillator 22A includesCsI (Tl) (cesium iodide having thallium added thereto) and thescintillator 22B includes gadolinium oxysulfide (GOS). In addition, acombination of the composition of the scintillator 22A and thecomposition of the scintillator 22B is not limited to theabove-mentioned example and may be a combination of other compositionsor a combination of the same compositions.

For example, the scintillators 22A and 22B have emission characteristicsthat vary depending on a thickness. As the thickness increases, theamount of light emitted increases and sensitivity increases. However,image quality deteriorates due to, for example, light scattering.

For example, in a case in which the scintillators 22A and 22B are formedby being filled with particles which are irradiated with the radiation Rand emit light, such as GOS particles, as the diameter of the particleincreases, the amount of light emitted increases and sensitivityincreases. However, the amount of light scattering increases and theincrease in the amount of light scattering affects adjacent pixels 32(see FIG. 3), which results in the deterioration of image quality.

In addition, the scintillators 22A and 22B may have a multi-layeredstructure of a small-particle layer and a large-particle layer. Forexample, in a case in which each of the first radiation detector 20A andthe second radiation detector 20B is irradiated with the radiation Rfrom the scintillators 22A and 22B to the TFT substrates 30A and 30Bunlike the radiography apparatus 16 according to this embodiment, thefollowing occurs. That is, in this case, image blurring is small in thescintillators 22A and 22B in which a region close to the irradiationside of the radiation R is filled with small particles and a regionclose to the side of the TFT substrate 30 that is the emission side ofthe radiation R is filled with large particles. However, obliquecomponents of light that is radially emitted by the small particles areless likely to reach the TFT substrates 30A and 30B and sensitivity isreduced. In addition, in a case in which the ratio of the region filledwith small particles to the region filled with large particles ischanged such that the number of layers formed by the region filled withlarge particles is larger than the number of layers formed by the regionfilled with small particles, sensitivity increases. However, in thiscase, light scattering affects adjacent pixels 32, which results in thedeterioration of image quality.

As the filling rate of the particles increases, the sensitivity of thescintillators 22A and 22B increases. However, the amount of lightscattering increases and image quality deteriorates. Here, the fillingrate is a value obtained by dividing the total volume of the particlesof the scintillator 22A or 22B by the volume of the scintillator 22A or22B and multiplying the divided value by 100 (the total volume of theparticles of the scintillator 22A or 22B/the volume of the scintillator22A or 22B×100). In addition, powder is treated in the scintillators 22Aand 22B. Therefore, in a case in which the filling rate is greater than80%, it is difficult to manufacture the scintillators 22A and 22B. Forthis reason, it is preferable that the filling rate is in the range of50 vol % to 80 vol %.

In addition, the emission characteristics of the scintillators 22A and22B vary depending on the doping amount of activator. As the dopingamount of activator increases, the amount of light emitted tends toincrease. However, the amount of light scattering increases and imagequality deteriorates.

The emission characteristics of the scintillators 22A and 22B withrespect to the radiation R vary depending on the material used for thescintillators 22A and 22B. For example, in a case in which each of thefirst radiation detector 20A and the second radiation detector 20B isirradiated with the radiation R from the scintillators 22A and 22B tothe TFT substrates 30A and 30B unlike the radiography apparatus 16according to this embodiment, the scintillator 22A is made of CsI (Tl)and the scintillator 22B is made of GOS. In this case, in thescintillator 22A, emphasis is put on image quality and the absorptivityof the low-energy radiation R is relatively high. In the scintillator22B, the absorptivity of the high-energy radiation R is relatively high.

In addition, the scintillator 22A has a columnar separated layerstructure, which makes it possible to further improve image quality.

In a case in which reflecting layers that transmit the radiation R andreflect visible light are formed on the surfaces of the scintillators22A and 22B which are opposite to the TFT substrates 30A and 30B, lightgenerated by the scintillators 22A and 22B is more effectively guided tothe TFT substrates 30A and 30B and sensitivity is improved. A method forforming the reflecting layer is not particularly limited. For example,any one of a sputtering method, a vapor deposition method, or a coatingmethod may be used. It is preferable that the reflecting layer is madeof a material with high reflectance in an emission wavelength range ofthe scintillators 22A and 22B used. For example, the reflecting layer ismade of Au, Ag, Cu, Al, Ni, and Ti. For example, in a case in which thescintillators 22A and 22B are made of GOS:Tb, the reflecting layer ispreferably made of Ag, Al, and Cu that have high reflectance in awavelength of 400 nm to 600 nm. In a case in which the thickness of thereflecting layer is less than 0.01 μm, reflectance is not obtained. Evenin a case in which the thickness is greater than 3 μm, the effect offurther improving the reflectance is not obtained. For this reason, itis preferable that the thickness of the reflecting layer is in the rangeof 0.01 μm to 3 μm.

Therefore, the characteristics of the scintillators 22A and 22B may varydepending on the diameter of particles, the multi-layered structure ofparticles, the filling rate of particles, the doping amount ofactivator, a material, a change in layer structure, and the shape of thereflecting layer.

In addition, the grid 23 that removes scattered rays generated by thetransmission of the radiation R through the subject W from the radiationR transmitted through the subject W is provided on the side of the firstradiation detector 20A on which the radiation R is incident side. Forexample, the effect of suppressing a reduction in the contrast of aradiographic image is obtained by the removal of the scattered rays fromthe radiation R and the quality of the radiographic image is improved.

The radiation limitation member 24 that limits the transmission of theradiation R is provided between the first radiation detector 20A and thesecond radiation detector 20B. An example of the radiation limitationmember 24 is a plate-shaped member made of, for example, copper or tin.It is preferable that the thickness of the plate-shaped member isuniform in the range in which an error of a variation in the thicknessis equal to or less than 1%. In a case in which the first radiationdetector 20A sufficiently absorbs the radiation R, the radiationlimitation member 24 may not be provided.

The control substrate 26A is provided so as to correspond to the firstradiation detector 20A and electronic circuits, such as an image memory56A and a control unit 58A which will be described below, are formed onthe control substrate 26A. The control substrate 26B is provided so asto correspond to the second radiation detector 20B and electroniccircuits, such as an image memory 56B and a control unit 58B which willbe described below, are formed on the control substrate 26B. The controlsubstrate 26A and the control substrate 26B are provided on the side ofthe second radiation detector 20B which is opposite to the incident sideof the radiation R.

The case 28 is provided at a position (that is, outside the range of animaging region) that does not overlap the radiation detector 20 at oneend of the housing 21. For example, a power supply unit 70 which will bedescribed below is accommodated in the case 28. The installationposition of the case 28 is not particularly limited. For example, thecase 28 may be provided at a position that overlaps the radiationdetector 20 on the side of the second radiation detector 20B which isopposite to the incident side of the radiation.

Next, the configuration of a main portion of an electric system of theradiography apparatus 16 according to this embodiment will be describedwith reference to FIG. 3.

As illustrated in FIG. 3, a plurality of pixels 32 are two-dimensionallyprovided in one direction (a row direction in FIG. 3) and a crossdirection (a column direction in FIG. 3) that intersects the onedirection on the TFT substrate 30A. The pixel 32 includes a sensor unit32A and a field effect thin film transistor (TFT; hereinafter, simplyreferred to as a “thin film transistor”) 32B.

The sensor unit 32A includes, for example, an upper electrode, a lowerelectrode, and a photoelectric conversion film which are notillustrated, absorbs the light emitted from the scintillator 22A,generates charge, and accumulates the generated charge. The thin filmtransistor 32B reads the charge accumulated in the sensor unit 32A,converts the charge into an electric signal, and outputs the electricsignal in response to a control signal. The sensor unit 32A is anexample of a conversion element that generates a larger amount of chargeas the amount of radiation becomes larger.

A plurality of gate lines 34 which extend in the one direction and areused to turn each thin film transistor 32B on and off are provided onthe TFT substrate 30A. In addition, a plurality of data lines 36 whichextend in the cross direction and are used to read out the chargethrough the thin film transistors 32B in an on state are provided on theTFT substrate 30A.

A gate line driver 52A is provided on one side of two adjacent sides ofthe TFT substrate 30A and a signal processing unit 54A is provided onthe other side. Each gate line 34 of the TFT substrate 30A is connectedto the gate line driver 52A and each data line 36 of the TFT substrate30A is connected to the signal processing unit 54A.

The rows of the thin film transistors 32B of the TFT substrate 30A aresequentially turned on by the electric signals which are supplied fromthe gate line driver 52A through the gate lines 34. Then, the chargewhich has been read out by the thin film transistor 32B in an on stateis transmitted as an electric signal through the data line 36 and isinput to the signal processing unit 54A. In this way, charge issequentially read out from each row of the thin film transistors andimage data indicating a two-dimensional radiographic image is acquired.

The signal processing unit 54A includes amplifying circuits (notillustrated) for amplifying an input electric signal and sample-and-holdcircuits (not illustrated) which are provided for each data line 36. Theelectric signal transmitted through each data line 36 is amplified bythe amplifying circuit and is then held by the sample-and-hold circuit.A multiplexer and an analog/digital (A/D) converter are connected to theoutput side of the sample-and-hold circuit in this order. The electricsignals held by each sample-and-hold circuit are sequentially (serially)input to the multiplexer and are sequentially selected by themultiplexer. Then, the selected electric signal is converted intodigital image data by the A/D converter.

The control unit 58A which will be described below is connected to thesignal processing unit 54A. The image data output from the A/D converterof the signal processing unit 54A is sequentially output to the controlunit 58A. The image memory 56A is connected to the control unit 58A. Theimage data sequentially output from the signal processing unit 54A issequentially stored in the image memory 56A under the control of thecontrol unit 58A. The image memory 56A has memory capacity that canstore a predetermined amount of image data. Whenever a radiographicimage is captured, captured image data is sequentially stored in theimage memory 56A.

The control unit 58A includes a central processing unit (CPU) 60, amemory 62 including, for example, a read only memory (ROM) and a randomaccess memory (RAM), and a non-volatile storage unit 64 such as a flashmemory. An example of the control unit 58A is a microcomputer.

A communication unit 66 is connected to the control unit 58A andtransmits and receives various kinds of information to and from externalapparatuses, such as the radiation emitting apparatus 12 and the console18, using at least one of wireless communication or wired communication.The power supply unit 70 supplies power to each of the above-mentionedvarious circuits or elements (for example, the gate line driver 52A, thesignal processing unit 54A, the image memory 56A, the control unit 58A,and the communication unit 66). In FIG. 3, lines for connecting thepower supply unit 70 to various circuits or elements are not illustratedin order to avoid complication.

Components of the TFT substrate 30B, the gate line driver 52B, thesignal processing unit 54B, the image memory 56B, and the control unit58B of the second radiation detector 20B have the same configurations asthe corresponding components of the first radiation detector 20A andthus the description thereof will not be repeated here. In addition, thecontrol unit 58A and the control unit 58B are connected such that theycan communicate with each other.

With the above-mentioned configuration, the radiography apparatus 16according to this embodiment captures radiographic images using thefirst radiation detector 20A and the second radiation detector 20B.Hereinafter, the radiographic image captured by the first radiationdetector 20A is referred to as a “first radiographic image” and imagedata indicating the first radiographic image is referred to as “firstradiographic image data”. In addition, hereinafter, the radiographicimage captured by the second radiation detector 20B is referred to as a“second radiographic image” and image data indicating the secondradiographic image is referred to as “second radiographic image data”.

Next, the configuration of the console 18 according to this embodimentwill be described with reference to FIG. 4. As illustrated in FIG. 4,the console 18 comprises a CPU 80 that controls the overall operation ofthe console 18 and a ROM 82 in which, for example, various programs orvarious parameters are stored in advance. In addition, the console 18comprises a RAM 84 that is used as, for example, a work area in a casein which the CPU 80 executes various programs and a non-volatile storageunit 86 such as a hard disk drive (HDD).

The console 18 further comprises a display unit 88 that displays, forexample, an operation menu and a captured radiographic image and anoperation panel 90 which includes a plurality of keys and to whichvarious kinds of information or operation commands are input. Inaddition, the console 18 comprises a communication unit 92 thattransmits and receives various kinds of information to and from theexternal apparatuses, such as the radiation emitting apparatus 12 andthe radiography apparatus 16, using at least one of wirelesscommunication or wired communication. The CPU 80, the ROM 82, the RAM84, the storage unit 86, the display unit 88, the operation panel 90,and the communication unit 92 are connected to each other through a bus94.

A plurality of correction data items 95 for correcting scattered raysare stored in the storage unit 86. In addition, the correction data 95will be described in detail below.

In the radiography apparatus 16 according to this embodiment, since thefirst radiation detector 20A and the radiation limitation member 24absorb the radiation R, the amount of radiation that reaches the secondradiation detector 20B is less than the amount of radiation that reachesthe first radiation detector 20A. In addition, the radiation limitationmember 24 generally has the characteristic that it absorbs a largernumber of soft-ray components than hard-ray components in energy formingthe radiation R, which depends on the material forming the radiationlimitation member 24. Therefore, the energy distribution of theradiation R that reaches the second radiation detector 20B has a largernumber of hard-ray components than the energy distribution of theradiation R that reaches the first radiation detector 20A.

In this embodiment, for example, about 50% of the radiation R that hasreached the first radiation detector 20A is absorbed by the firstradiation detector 20A and is used to capture a radiographic image. Inaddition, about 60% of the radiation R that has passed through the firstradiation detector 20A and reached the radiation limitation member 24 isabsorbed by the radiation limitation member 24. About 50% of theradiation R that has passed through the first radiation detector 20A andthe radiation limitation member 24 and reached the second radiationdetector 20B is absorbed by the second radiation detector 20B and isused to capture a radiographic image. Since the absorptivity ofradiation by the radiation detector 20 and the radiation limitationmember 24 varies depending on the energy of the radiation R, the shapeof a spectrum changes.

That is, the amount of radiation used by the second radiation detector20B to capture a radiographic image is about 20% of the amount ofradiation used by the first radiation detector 20A to capture aradiographic image. In addition, the ratio of the amount of radiationused by the second radiation detector 20B to capture a radiographicimage to the amount of radiation used by the first radiation detector20A to capture a radiographic image is not limited to theabove-mentioned ratio. However, it is preferable that the amount ofradiation used by the second radiation detector 20B to capture aradiographic image is equal to or greater than 10% of the amount ofradiation used by the first radiation detector 20A to capture aradiographic image in terms of diagnosis.

Low-energy components of the radiation R are absorbed first. Theradiation R absorbed by each of the first radiation detector 20A and thesecond radiation detector 20B will be described with reference to FIG.5. In FIG. 5, the vertical axis indicates the amount of radiation Rabsorbed and the horizontal axis indicates the energy of the radiation Rin a case in which the tube voltage of the radiation source 14 is 80 kV.In addition, in FIG. 5, a solid line L1 indicates the relationshipbetween the energy of the radiation R absorbed by the first radiationdetector 20A and the amount of radiation R absorbed. In addition, inFIG. 5, a solid line L2 indicates the relationship between the energy ofthe radiation R absorbed by the second radiation detector 20B and theamount of radiation R absorbed. Since the low-energy components of theradiation R are absorbed first, for example, as illustrated in FIG. 5,the energy components of the radiation R that reaches the secondradiation detector 20B do not include the low-energy components of theenergy components of the radiation R that reaches the first radiationdetector 20A. That is, the energy of the radiation R emitted to thefirst radiation detector 20A is different from the energy of theradiation R emitted to the second radiation detector 20B through thefirst radiation detector 20A. Therefore, in the radiography apparatus 16according to this embodiment, the radiation detectors 20 are irradiatedwith the radiations R having different energy levels (radiation R with afirst energy level and radiation R with a second energy level) andradiographic images are generated by the radiation detectors 20.

The console 18 according to this embodiment acquires radiographic imagedata generated by the radiation detectors 20 irradiated with theradiations R having different energy levels. In addition, the console 18derives the ratio of the values of the corresponding pixels of firstradiographic image data and second radiographic image data and generatesimage data for deriving the bone density of the subject W. Hereinafter,the image data for deriving the bone density of the subject W isreferred to as “DXA image data” and an image indicated by the DXA imagedata is referred to as a “DXA image”. Specifically, the console 18performs log conversion for each pixel value of each of the firstradiographic image data and the second radiographic image data. Then,the console 18 generates DXA image data, using an energy subtractionprocess that subtracts image data obtained by performing log conversionfor the second radiographic image data from image data obtained byperforming log conversion for the first radiographic image data for eachcorresponding pixel.

In addition, for example, as illustrated in FIG. 6, the console 18according to this embodiment derives bone density from each pixel value(that is, the ratio of the values of the corresponding pixels of thefirst radiographic image and the second radiographic image and adifference value between the pixel values in a log image) of a bone partof the subject W in the cross-sectional direction (the horizontaldirection in the example illustrated in FIG. 6) in the DXA image.

FIG. 7 illustrates the value of each pixel in a region R1 of the DXAimage illustrated in FIG. 6. In FIG. 7, the horizontal axis indicates apixel position in the horizontal direction of FIG. 6. In addition, inFIG. 7, the vertical axis indicates an average value of the values of aplurality of pixels in the vertical direction of FIG. 6 at each pixelposition in the horizontal direction of FIG. 6. Hereinafter, a datagroup of the pixel values at each pixel position along the horizontaldirection of FIG. 6 which is illustrated in FIG. 7 is referred to as a“DXA profile”.

As illustrated in FIG. 7, for the pixel values in the DXA profile, apixel value at a pixel position corresponding to the bone tissue of thesubject W is less than a pixel value at a pixel position correspondingto the soft tissue. The console 18 according to this embodiment derivesthe average value of the pixel values in soft tissue regions(hereinafter, referred to as “soft regions”) on both sides of a bonetissue region (hereinafter, referred to as a “bone region”) and derivesa straight line (hereinafter, referred to as a “reference line”) K thatconnects the average values derived at the pixel positions in a centralportion of each soft region. In addition, the console 18 adds thedifferences between the reference line K and the pixel values at eachpixel position in the bone region to derive the area of the bone region(the area of a hatched portion illustrated in FIG. 7). The area is avalue corresponding to the bone mass of the subject W.

In addition, the console 18 divides the derived area by the number ofpixels corresponding to the width of the bone region to derive thedifference between the pixel values of the bone region and the softregion per unit number of pixels. The difference is a valuecorresponding to the bone density of the subject W. Then, the console 18multiplies the derived difference between the pixel values of the boneregion and the soft region per unit number of pixels by a predeterminedunit conversion coefficient to derive the bone density of the subject W.In this embodiment, the pixel position of the region R1 used to derivethe DXA profile in the DXA image data, the pixel position of the softregion of the DXA profile, and the pixel position of the bone region arepredetermined according to, for example, the subject W and an imagingpart.

A predetermined amount of scattered rays is removed by the grid 23.However, components (hereinafter, “scattered ray components”) caused bythe scattered rays which have not been removed by the grid 23 areincluded in the first radiographic image and the second radiographicimage. In particular, in a DXA method for deriving the bone density ofthe subject W, the numerical value of the derived bone density is alsoaffected by the amount of scattered rays which have not been removed bythe grid 23. For this reason, the console 18 according to thisembodiment corrects the scattered ray components included in the firstradiographic image and the second radiographic image, using thecorrection data 95. In addition, the intensity and spread of thescattered rays vary depending on various conditions. Therefore, in thisembodiment, calibration is performed according to various conditions andthe correction data 95 obtained by the calibration is stored in thestorage unit 86 so as to be associated with each combination of theconditions. Hereinafter, the calibration will be described in detail.

The correction data 95 according to this embodiment will be describedwith reference to FIGS. 8A to 10. In this embodiment, as illustrated inFIG. 8A, a plurality of correction data items 95 which have beenobtained in advance by calibration using a phantom PT simulating thehuman body and a flat-plate-shaped radiation shielding member Bshielding the radiation R are stored in the storage unit 86. A pinholePH is formed in a central portion of the radiation shielding member B.The phantom PT simulates the human body using a material correspondingto the soft tissues of the human body and a material corresponding tothe bone tissues of the human body. For example, acryl or urethane canbe applied as the material corresponding to the soft tissues of thehuman body. In addition, for example, hydroxyapatite can be applied asthe material corresponding to the bone tissues of the human body.Hereinafter, the correction data 95 will be described in detail.

In each radiation detector 20, different scattered rays are generated ina region (hereinafter, referred to as a “subject region”) irradiatedwith the radiation R that has been transmitted through the subject W anda region (hereinafter, referred to as a “directly irradiated region”)directly irradiated with the radiation R that has not been transmittedthrough the subject W. In addition, in the subject region, differentscattered rays are generated in parts of the subject W, such as softtissues, bone tissues, and the lungs. Therefore, in this embodiment, forcalibration related to the subject region, as illustrated in FIG. 8B,the phantom PT is disposed on the side of the radiography apparatus 16on which the radiation R is incident, the radiation shielding member Bis disposed on the side of the phantom PT on which the radiation R isincident, and the radiation R is emitted from the radiation emittingapparatus 12 for a predetermined period. The console 18 generates thecorrection data 95 with image data obtained from each radiation detector20 in this case and stores the generate correction data 95 as datacorresponding to the subject region and each radiation detector 20 inthe storage unit 86.

In this embodiment, in the case of the calibration related to thesubject region, the correction data 95 is generated for each region ofthe radiographic image which corresponds to each part of the subject Wwith the phantom PT using the materials corresponding to each part.Then, the generated correction data 95 is stored in the storage unit 86so as to be associated with each region.

For example, the correction data 95 corresponding to each regionincludes information indicating the spread of scattered rays illustratedin FIG. 9A and information indicating the intensity of scattered raysillustrated in FIG. 9B. The information indicating the spread ofscattered rays is also referred to as a point spread function (PSF). Inaddition, the information indicating the intensity of scattered rays isinformation in which a pixel value per unit amount of radiation isassociated with the ratio of scattered rays at the pixel value.

As described above, the energy of the radiation R emitted to the firstradiation detector 20A is different from the energy of the radiation Remitted to the second radiation detector 20B. Therefore, as illustratedin FIGS. 9A and 9B, the information indicating the spread of scatteredrays and the information indicating the intensity of scattered rays aredifferent in the first radiation detector 20A and the second radiationdetector 20B. For this reason, in this embodiment, the correction data95 obtained from the first radiation detector 20A by calibration isstored in the storage unit 86 so as to be associated with the firstradiation detector 20A. In addition, the correction data 95 obtainedfrom the second radiation detector 20B by calibration is stored in thestorage unit 86 so as to be associated with the second radiationdetector 20B.

In a soft region of the subject region, different scattered rays aregenerated according to the body thickness of the subject W. Therefore,in this embodiment, as illustrated in FIG. 8B, for the soft region, thecorrection data 95 obtained by calibration using the phantoms PT with aplurality of types of body thicknesses is stored in the storage unit 86so as to be associated with each body thickness.

In contrast, for calibration related to the directly irradiated region,as illustrated in FIG. 8C, the radiation shielding member B is disposedon the side of the radiography apparatus 16 on which the radiation R isincident and the radiation R is emitted from the radiation emittingapparatus 12 for a predetermined period. The console 18 derivesinformation indicating the spread of scattered rays, using image dataobtained from each radiation detector 20 in this case. In addition, theconsole 18 derives the amount of scattered rays from the amount ofradiation and derives the amount of scattered rays per unit amount ofradiation as information indicating the intensity of scattered rays.Then, the console 18 stores the derived information indicating thespread of scattered rays and the derived information indicating theintensity of scattered rays as the correction data 95 corresponding tothe directly irradiated region and each radiation detector 20 in thestorage unit 86.

In addition, different scattered rays are generated according to imagingconditions. Therefore, in this embodiment, calibration is performed eachof the imaging conditions used in the facility in which the radiographysystem 10 is provided and the correction data 95 is stored in thestorage unit 86 so as to be associated with the imaging conditions. Theimaging conditions include, for example, a material (for example,tungsten) forming a bulb of the radiation source 14, a tube voltage, amaterial (for example, copper) forming the radiation limitation member24, the characteristics of the grid 23 (for example, a grid ratio, griddensity, and a convergence distance), and a source image distance (SID).The SID indicates a distance from the radiation source 14 to a surfacefor detecting the radiation R in the first radiation detector 20A. Inaddition, the imaging conditions include the quality of a material (forexample, carbon) forming the surface of a case accommodating theradiography apparatus 16 on which the radiation R is incident. Examplesof the surface of the case accommodating the radiography apparatus 16 onwhich the radiation R is incident include a top plate of a decubitusimaging table and a decorative cover of an upright imaging table.

As illustrated in FIG. 10, in a case in which a radiographic image iscaptured, a gap is formed between a surface of the subject W which facesthe radiography apparatus 16 and a surface S of the case on which theradiation R is incident. In addition, as the distance d of the gapbecomes longer, the scattered rays are more widely spread and theintensity of the scattered rays becomes lower. Therefore, in thisembodiment, calibration is performed for each of a plurality ofdistances d and the correction data 95 is stored in the storage unit 86so as to be associated with each of the plurality of distances d foreach radiation detector 20. In this embodiment, the distance d of thegap means a distance between the surface of the subject W which facesthe radiography apparatus 16 and the surface S of the case on which theradiation R is incident in the direction of the body thickness.

Next, the operation of the radiography system 10 according to thisembodiment will be described with reference to FIGS. 11 to 13. FIG. 11is a flowchart illustrating the process flow of an overall imagingprocessing program executed by the CPU 80 of the console 18 in a case inwhich the user inputs the name of the subject W, an imaging part, and animaging menu through the operation panel 90. The overall imagingprocessing program is installed in the storage unit 86 of the console 18in advance.

FIG. 12 is a flowchart illustrating the process flow of an individualimaging processing program executed by the control unit 58A of theradiography apparatus 16 in a case in which the radiography apparatus 16is turned on. The individual imaging processing program is installed inthe ROM of the memory 62 of the control unit 58A in advance. Inaddition, the individual imaging processing program is installed in theROM of the memory 62 of the control unit 58B in advance and is executedby the control unit 58B of the radiography apparatus 16 in a case inwhich the radiography apparatus 16 is turned on. In the individualimaging process illustrated in FIG. 12, the control unit 58A and thecontrol unit 58B perform the same process. Therefore, hereinafter, onlya case in which the individual imaging process is performed by thecontrol unit 58A will be described and the description of a case inwhich the individual imaging process is performed by the control unit58B will be omitted.

In Step 100 illustrated in FIG. 11, the CPU 80 transmits informationincluded in the input imaging menu to the radiography apparatus 16through the communication unit 92 and transmits the emission conditionsof the radiation R to the radiation emitting apparatus 12 through thecommunication unit 92. Then, the CPU 80 transmits a command to start theemission of the radiation R to the radiography apparatus 16 and theradiation emitting apparatus 12 through the communication unit 92. In acase in which the emission conditions and the emission start commandtransmitted from the console 18 are received, the radiation emittingapparatus 12 starts the emission of the radiation R according to thereceived emission conditions. The radiation emitting apparatus 12 mayinclude an irradiation button. In this case, the radiation emittingapparatus 12 receives the emission conditions and the emission startcommand transmitted from the console 18 and starts the emission of theradiation R according to the received emission conditions in a case inwhich the irradiation button is pressed.

In Step 102, the CPU 80 waits until the first radiographic image datacaptured by the first radiation detector 20A and the second radiographicimage data captured by the second radiation detector 20B are received.In a case in which the CPU 80 receives the first radiographic image dataand the second radiographic image data through the communication unit92, the determination result in Step 102 is “Yes” and the processproceeds to Step 104.

In Step 104, the CPU 80 performs a bone density derivation processillustrated in FIG. 13 and then ends the overall imaging process.

In Step 120 of FIG. 12, the control unit 58A performs a reset operationwhich extracts the charge accumulated in the sensor unit 32A of eachpixel 32 in the first radiation detector 20A and removes the charge. Inaddition, the control unit 58A may perform the reset operation in Step120 only once, may repeatedly perform the reset operation apredetermined number of times, or may repeatedly perform the resetoperation until the determination result in Step 122, which will bedescribed below, becomes “Yes”.

In Step 122, the control unit 58A waits until a command to start theemission of the radiation R is received. In a case in which the controlunit 58A receives the emission start command transmitted from theconsole 18 in Step 100 of the overall imaging process through thecommunication unit 66, the determination result in Step 122 is “Yes” andthe process proceeds to Step 124. In a case in which the radiationemitting apparatus 12 comprises an irradiation button and the controlunit 58A receives the emission start command transmitted from theconsole 18 and information indicating that the irradiation button hasbeen pressed through the communication unit 66, the determination resultin Step 122 is “Yes”. For example, in a case in which the irradiationbutton is pressed, the radiation emitting apparatus 12 may directlytransmit information indicating that the irradiation button has beenpressed to the radiography apparatus 16 or may transmit the informationto the radiography apparatus 16 through the console 18.

In Step 124, the control unit 58A waits for an emission period that isincluded in the information transmitted from the console 18 in Step 100of the overall imaging process.

In Step 126, the control unit 58A controls the gate line driver 52A suchthat the gate line driver 52A sequentially outputs an on signal to eachof the gate lines 34 of the first radiation detector 20A for apredetermined period. Then, the rows of the thin film transistors 32Bconnected to each gate line 34 are sequentially turned on and the chargeaccumulated in each sensor unit 32A in each row sequentially flows as anelectric signal to each data line 36. Then, the electric signal whichhas flowed to each data line 36 is converted into digital image data bythe signal processing unit 54A and is stored in the image memory 56A.

In Step 128, the control unit 58A performs image processing forperforming various types of correction, such as offset correction andgain correction, for the image data stored in the image memory 56A inStep 126. In Step 130, the control unit 58A transmits the image data(first radiographic image data) subjected to the image processing inStep 128 to the console 18 through the communication unit 66 and thenends the individual imaging process.

In a case in which the console 18 receives the first radiographic imagedata and the second radiographic image data transmitted in Step 130, thedetermination result in Step 102 is “Yes” and the bone densityderivation process illustrated in FIG. 13 is performed.

In Step 140 of FIG. 13, the CPU 80 stores the first radiographic imagedata and the second radiographic image data received in Step 102 in thestorage unit 86. In Step 142, the CPU 80 specifies a region(hereinafter, a “radiation shielded region”) in which the radiation R isshielded, the directly irradiated region, and the subject region in thefirst radiographic image, using the first radiographic image datareceived in Step 102. The radiation shielded region corresponds to, forexample, a region which is not irradiated with the radiation R by acollimator.

Specifically, the CPU 80 specifies, as the radiation shielded region, aregion having a pixel value that is equal to or less than a pixel valueobtained by adding a predetermined margin to a pixel value predeterminedas the pixel value of the region that is not irradiated with theradiation R in the first radiographic image data. In addition, in a casein which configuration information including the distance and thepositional relationship among the radiation source 14, the collimator,the radiography apparatus 16 and the size of the collimator can beacquired, the console 18 may specify the radiation shielded region fromthe configuration information.

In addition, the CPU 80 specifies, as the directly irradiated region, aregion in which a pixel value is saturated in the first radiographicimage data. Further, the CPU 80 specifies, as the subject region, aregion other than the radiation shielded region and the directlyirradiated region in the first radiographic image. Furthermore, the CPU80 performs an image analysis process for a portion of the subjectregion in the first radiographic image to specify regions of the partsof the subject W corresponding to imaging parts, such as a soft region,a bone region, and a lung region. In addition, similarly to the firstradiographic image, the CPU 80 specifies each region of the parts of thesubject W in the radiation shielded region, the directly irradiatedregion, the subject region, and the subject region of the secondradiographic image, using the second radiographic image data received inStep 102. For example, the user may input information indicating eachregion through the operation panel 90.

In Step 144, for example, as illustrated in FIG. 14, the CPU 80generates an image (hereinafter, referred to as a “directly irradiatedregion image”) by extracting the directly irradiated region from thefirst radiographic image. In addition, the CPU 80 acquires thecorrection data 95 associated with the imaging conditions, the directlyirradiated region, and the first radiation detector 20A from the storageunit 86. Then, the CPU 80 generates a scattered ray image caused by thedirectly irradiated region, using the directly irradiated region imageand the acquired correction data 95. Specifically, the CPU 80 derivesthe amount and spread of scattered rays for each pixel of the directlyirradiated region of the directly irradiated region image, using thecorrection data 95, and performs a convolution operation for the derivedamount and spread of scattered rays to generate the scattered ray imagecaused by the directly irradiated region. In addition, for example, theconsole 18 may acquire the correction data 95 from an external systemconnected through the network.

Further, similarly, for the second radiographic image, the CPU 80generates a scattered ray image caused by the directly irradiatedregion, using the directly irradiated region image and the correctiondata 95 associated with the imaging conditions, the directly irradiatedregion, and the second radiation detector 20B.

In Step 146, the CPU 80 generates a scattered ray image caused by eachregion of each part of the subject W in the subject region specified inStep 142. For example, for the soft region, as illustrated in FIG. 15,the CPU 80 generates an image (hereinafter, referred to as a “bodythickness image”) in which each pixel has a pixel value corresponding tothe body thickness of the subject W from the first radiographic image. Alarger amount of radiation R is absorbed as the body thickness of thesubject W becomes larger. Therefore, in many cases, the pixel value ofthe soft region becomes smaller as the body thickness of the subject Wbecomes larger. For this reason, for example, the CPU 80 estimates thebody thickness of the subject W from the pixel value of the soft region,using information in which the body thickness is associated with thepixel value of the soft region, and generates a body thickness imageaccording to the estimated body thickness.

Then, the CPU 80 performs the convolution operation for each pixel ofthe body thickness image, using the correction data 95 associated withthe imaging conditions, the body thickness, and the first radiationdetector 20A, to generate a scattered ray image caused by the softregion as in Step 144. Similarly, for the second radiographic image, theCPU 80 performs the convolution operation for each pixel of the bodythickness image, using the correction data 95 associated with theimaging conditions, the body thickness, and the second radiationdetector 20B, to generate a scattered ray image caused by the softregion.

In addition, for each region of the subject region other than the softregion, the CPU 80 generates each image by extracting each region of thefirst radiographic image as in Step 144. Then, the CPU 80 generates eachscattered ray image caused by each region, using the generated image ofeach region and the correction data 95 associated with the imagingconditions, each corresponding region, and the first radiation detector20A. Similarly, for the second radiographic image, the CPU 80 generateseach scattered ray image caused by each region, using the image of eachregion generated from the second radiographic image and the correctiondata 95 associated with the imaging conditions, each correspondingregion, and the second radiation detector 20B.

In Step 148, the CPU 80 derives the distance d using the firstradiographic image. A process of deriving the distance d will bedescribed in detail with reference to FIGS. 16 to 18.

The CPU 80 may generate an energy subtraction image (hereinafter,referred to as an “ES image”) using the first radiographic image dataand the second radiographic image data. Specifically, the CPU 80subtracts image data obtained by multiplying the first radiographicimage data by a predetermined coefficient from image data obtained bymultiplying the second radiographic image data by a predeterminedcoefficient for each corresponding pixel. The CPU 80 generates an ESimage in which the bone tissues have been removed and the soft tissueshave been highlighted, using the subtraction.

In addition, the CPU 80 applies a low-pass filter for smoothing theinfluence of each part of the subject W to the generated ES image togenerate an image (hereinafter, referred to as a gap deriving image”)for deriving a gap. Further, for example, as illustrated in FIG. 16, theCPU 80 derives the average value of the values of a plurality of pixelsin a direction perpendicular to the cross-sectional direction at eachpixel position in the cross-sectional direction of a bone part in aregion, from which bone density is to be derived, in the gap derivingimage, similarly to the DXA profile. Hereinafter, a data group of thepixel values illustrated in FIG. 16 is referred to as a “gap derivingprofile”.

As illustrated in FIG. 16, the gap deriving profile is substantiallyflat at a pixel position corresponding to a portion of the subject Wwhich comes into contact with the surface S and the pixel value becomeslarger as the distance d becomes longer. As illustrated in FIG. 17, thepixel value has the property that it is exponentially reduced as thebody thickness increases. Therefore, as illustrated in FIG. 18, the CPU80 can derive a first body thickness d1 of a central portion of thesubject W from a pixel value QL1 of the center of a portion in which theshape of the gap deriving profile is substantially flat. In addition,the CPU 80 can derive a second body thickness d2 of a portion in whichthere is a gap between the surface of the subject W and the surface S inthe subject W from a pixel value QL2 at a pixel position where the pixelvalue of the gap deriving profile increases gradually.

As illustrated in FIG. 18, the CPU 80 derives the distance d from thefirst body thickness d1 and the second body thickness d2 according tothe following Expression (1), assuming that the body thickness of thesubject W is symmetrically reduced from a side opposite to the surfaceS:

d=(d1−d2)÷2  (1).

In addition, the CPU 80 derives the distance d for each pixel positionwhere the pixel value of the gap deriving profile increases gradually.

In Step 150, the CPU 80 performs the convolution operation for eachpixel of a region corresponding to the pixel position where the distanced has been derived in the first radiographic image, using the correctiondata 95 associated with the imaging conditions, the distance d, and thefirst radiation detector 20A, to generate a scattered ray image causedby the distance d. In addition, similarly, for the second radiographicimage, the CPU 80 performs the convolution operation, using thecorrection data 95 associated with the imaging conditions, the distanced, and the second radiation detector 20B, to generate a scattered rayimage caused by the distance d.

In Step 152, the CPU 80 subtracts image data indicating each scatteredray image generated from the first radiographic image in Steps 144, 146,and 150 from the first radiographic image data for each correspondingpixel to correct the scattered ray components of the first radiographicimage. In addition, the CPU 80 subtracts image data indicating eachscattered ray image generated from the second radiographic image inSteps 144, 146, and 150 from the second radiographic image data for eachcorresponding pixel to correct the scattered ray components of thesecond radiographic image.

In Step 154, the CPU 80 generates DXA image data, using the correctedfirst radiographic image data and second radiographic image datasubjected to the correction in Step 152. In a case in which the firstradiographic image data and the second radiographic image data aresimply referred to in Steps 154 to 160, it is assumed that the firstradiographic image data and the second radiographic image data indicatethe corrected first radiographic image data and second radiographicimage data subjected to the correction in Step 152, respectively.

The CPU 80 performs log conversion for each pixel value of each of thefirst radiographic image data and the second radiographic image data.Then, the CPU 80 generates DXA image data, using an energy subtractionprocess that subtracts image data obtained by performing log conversionfor the second radiographic image data from image data obtained byperforming log conversion for the first radiographic image data for eachcorresponding pixel. Then, the CPU 80 stores the generated DXA imagedata in the storage unit 86.

A method for determining the corresponding pixels of the firstradiographic image data and the second radiographic image data is notparticularly limited. For example, the amount of positional deviationbetween first radiographic image data and second radiographic image datais calculated from a difference in the position of a marker between thefirst radiographic image data and the second radiographic image datacaptured by the radiography apparatus 16 in a state in which the markeris placed in advance. Then, the corresponding pixels of the firstradiographic image data and the second radiographic image data aredetermined on the basis of the calculated amount of positionaldeviation.

In this case, for example, the amount of positional deviation betweenfirst radiographic image data and second radiographic image data may becalculated from a difference in the position of a marker between thefirst radiographic image data and the second radiographic image dataobtained by capturing the image of the marker together with the subjectW in a case in which the image of the subject is captured. In addition,for example, the amount of positional deviation between firstradiographic image data and second radiographic image data may becalculated on the basis of the structure of the subject in the firstradiographic image data and the second radiographic image data obtainedby capturing the image of the subject.

In Step 156, the CPU 80 derives a DXA profile using the DXA image datagenerated in Step 154. In Step 158, the CPU 80 derives an integratedvalue of the differences between the reference line K and the pixelvalues of the bone region in the DXA profile derived in Step 156. Inaddition, the CPU 80 divides the derived integrated value by the numberof pixels corresponding to the width of the bone region in the DXAprofile. Then, the CPU 80 multiplies the value obtained by the divisionby a unit conversion coefficient to derive the bone density of thesubject W.

In Step 160, the CPU 80 displays a first radiographic image indicated bythe first radiographic image data as an image for diagnosis on thedisplay unit 88 and displays the bone density derived in Step 158 on thedisplay unit 88. Then, the CPU 80 ends the bone density derivationprocess.

In addition, the CPU 80 may generate an ES image, using the correctedfirst radiographic image data and second radiographic image datasubjected to the correction in Step 152. In this case, for example, theCPU 80 subtracts image data obtained by multiplying the firstradiographic image data by a predetermined coefficient from image dataobtained by multiplying the second radiographic image data by apredetermined coefficient for each corresponding pixel. The CPU 80generates an ES image in which the soft tissues have been removed andthe bone tissues have been highlighted, using the subtraction. In thisexample, in Step 158, the CPU 80 may display an ES image in which thebone tissues have been highlighted on the display unit 88, instead ofthe image for diagnosis.

In addition, the CPU 80 may specify the edge of a bone region from theES image in which the bone tissues have been highlighted and may use thespecification result as a pixel position corresponding to the boneregion in the DXA image data. In this case, for example, the CPU 80estimates the approximate range of the bone region on the basis of theimaging part included in the imaging menu. Then, the CPU 80 detectspixels that are disposed in the vicinity of the pixels, of which thedifferential values are equal to or greater than a predetermined value,as the pixels forming the edge (end) of the bone region in the estimatedrange to specify the bone region.

In this case, the CPU 80 may specify, as the soft region, a region whichhas a predetermined area including pixels that are separated from thespecified edge of the bone region by a distance corresponding to apredetermined number of pixels in a predetermined direction in which theregion becomes further away from the bone part. In this case, the CPU 80may use the specification result as a pixel position corresponding tothe soft tissue in the DXA image data.

As described above, according to this embodiment, scattered raycomponents included in a radiographic image are corrected using thecorrection data 95 for correcting scattered rays which is associatedwith each of a plurality of regions in the radiographic image.Therefore, it is possible to correct components caused by the scatteredrays of the radiation R included in the radiographic image with highaccuracy.

In addition, according to this embodiment, the bone density of thesubject W is derived using the radiographic images generated by each oftwo radiation detectors 20 provided in the radiography apparatus 16.Therefore, the bone density of the subject W can be derived by oneoperation of emitting the radiation R. As a result, the amount ofradiation R emitted to the subject W is reduced and it is possible toderive the bone density of the subject W.

Second Embodiment

Hereinafter, a second embodiment of the present disclosure will bedescribed in detail. Since the configuration of a radiography system 10according to this embodiment is the same as that in the first embodiment(see FIG. 1, FIG. 3, and FIG. 4) except the configuration of aradiography apparatus 16, the description thereof will not be repeatedhere. In addition, components having the same functions as those in thefirst embodiment are denoted by the same reference numerals and thedescription thereof will not be repeated.

As illustrated in FIG. 19, a radiation detector 20C that detectsradiation R transmitted through a subject W and a control substrate 26Care provided in the housing 21 of the radiography apparatus 16 accordingto this embodiment. Since the configuration of the radiation detector20C is the same as that of the first radiation detector 20A according tothe first embodiment, the description thereof will not be repeated here.In addition, since the configuration of the control substrate 26C is thesame as that of the control substrate 26A according to the firstembodiment, the description thereof will not be repeated here. A grid 23for removing scattered rays is provided between the housing 21 and thesubject W.

The radiography system 10 according to this embodiment performs tworadiography operations with different tube voltages and derives bonedensity, using radiographic image data items obtained from the radiationdetector 20C by two imaging operations. Since the tube voltages aredifferent in the two imaging operations, the radiation detector 20C isirradiated with the radiations R with different energy levels. Theradiation R absorbed by the radiation detector 20C will be describedwith reference to FIG. 20. In FIG. 20, the vertical axis indicates theamount of radiation R absorbed and the horizontal axis indicates theenergy of the radiation R. In addition, in FIG. 20, a solid line L3indicates the relationship between the energy of the radiation Rabsorbed by the radiation detector 20C and the amount of radiation Rabsorbed in a case in which the tube voltage of the radiation source 14is 140 kV. In FIG. 20, a solid line L4 indicates the relationshipbetween the energy of the radiation R absorbed by the radiation detector20C and the amount of radiation R absorbed in a case in which the tubevoltage of the radiation source 14 is 100 kV. As illustrated in FIG. 20,since the tube voltage of the radiation source 14 is different, theradiation detector 20C is irradiated with the radiations R withdifferent energy levels in first irradiation and second irradiation.

Correction data 95 according to this embodiment is stored in the storageunit 86 so as to be associated with a plurality of imaging conditionshaving at least different tube voltages. In other words, the correctiondata 95 is stored in the storage unit 86 so as to be associated with aplurality of different energy levels corresponding to the tube voltages.In addition, the correction data 95 according to this embodiment isobtained by the same calibration as that in the first embodiment.

Next, the operation of the radiography system 10 according to thisembodiment will be described with reference to FIGS. 21 to 23. In FIG.21, steps in which the same processes as those in FIG. 11 are performedare denoted by the same reference numerals as those in FIG. 11 and thedescription thereof will not be repeated. In FIG. 22, steps in which thesame processes as those in FIG. 12 are performed are denoted by the samereference numerals as those in FIG. 12 and the description thereof willnot be repeated. In FIG. 23, steps in which the same processes as thosein FIG. 13 are performed are denoted by the same reference numerals asthose in FIG. 13 and the description thereof will not be repeated.

In Step 103 of FIG. 21, the CPU 80 waits until radiographic image datacaptured by the radiation detector 20C is received. In a case in whichthe CPU 80 receives radiographic image data through the communicationunit 92, the determination result in Step 103 is “Yes” and the processproceeds to Step 105. In Step 105, the CPU 80 stores the radiographicimage data received in Step 103 in the storage unit 86 and then ends theoverall imaging process.

In this embodiment, the user performs the overall imaging process twotimes in a series of radiography processes. In this case, the user setsthe tube voltage so as to be different in a first imaging operation anda second imaging operation. In this embodiment, a case in which the tubevoltage (for example, 70 [kV]) in the first imaging operation is lowerthan the tube voltage in the second imaging operation and the tubevoltage (for example, 100 [kV]) in the second imaging operation ishigher than the tube voltage in the first imaging operation will bedescribed. In addition, the tube voltage in the first imaging operationmay be higher than the tube voltage in the second imaging operation.

In Step 131 of FIG. 22, the control unit 58A determines whether theprocess from Step 120 to Step 130 has been repeatedly performed twotimes. In a case in which the determination result is “No”, the processreturns to Step 120. In a case in which the determination result is“Yes”, the individual imaging process ends.

The CPU 80 of the console 18 performs the following process, using theradiographic image data transmitted from the radiography apparatus 16 bythe first process in Step 130 as the second radiographic image data.That is, in this case, the CPU 80 performs a bone density derivationprocess illustrated in FIG. 23, using the radiographic image datatransmitted from the radiography apparatus 16 by the second process inStep 130 as the first radiographic image data.

In Step 144A of FIG. 23, the CPU 80 generates a directly irradiatedregion image from the first radiographic image as in Step 144 accordingto the first embodiment. In addition, the CPU 80 acquires the correctiondata 95 associated with the imaging conditions of the second imagingoperation and the directly irradiated region from the storage unit 86.Then, the CPU 80 generates a scattered ray image caused by the directlyirradiated region, using the directly irradiated region image and theacquired correction data 95. In addition, similarly, for the secondradiographic image, the CPU 80 generates a scattered ray image caused bythe directly irradiated region, using the directly irradiated regionimage and the acquired correction data 95 associated with the imagingconditions of the first imaging operation and the directly irradiatedregion.

In Step 146A, the CPU 80 generates a scattered ray image caused by eachregion of each part of the subject W in the subject region specified inStep 142. For example, the CPU 80 generates a body thickness image fromthe first radiographic image as in Step 146 according to the firstembodiment. Then, the CPU 80 performs a convolution operation for eachpixel of the body thickness image, using the correction data 95associated with the imaging conditions of the second imaging operationand the body thickness, to generate a scattered ray image caused by thesoft region as in Step 144A. Similarly, for the second radiographicimage, the CPU 80 performs the convolution operation for each pixel ofan image indicated by the body thickness image, using the correctiondata 95 associated with the imaging conditions of the first imagingoperation and the body thickness, to generate a scattered ray imagecaused by the soft region.

In addition, for each region other than the soft region of the subjectregion, the CPU 80 generates an image by extracting each region from thefirst radiographic image as in Step 144A. Then, the CPU 80 generateseach scattered ray image caused by each region, using the generatedimage of each region and the correction data 95 associated with theimaging conditions of the second imaging operation and eachcorresponding region. Similarly, the CPU 80 generates each scattered rayimage caused by each region, using the image of each region generatedfrom the second radiographic image and the correction data 95 associatedwith the imaging conditions of the first imaging operation and eachcorresponding region.

In Step 150A, the CPU 80 performs the convolution operation for eachpixel of a region corresponding to the pixel position where the distanced is derived in the first radiographic image, using the correction data95 associated with the imaging conditions of the second imagingoperation and the distance d, to generate a scattered ray image causedby the distance d. Similarly, for the second radiographic image, the CPU80 performs the convolution operation, using the correction data 95associated with the imaging conditions of the first imaging operationand the distance d, to generate a scattered ray image caused by thedistance d.

In this embodiment, the same radiography apparatus 16 as that accordingto the first embodiment may be used. In this case, for example, a DXAimage is generated from each of the radiographic images generated byirradiating the first radiation detector 20A provided on the incidentside of the radiation R with the radiations R with different energylevels.

As described above, according to this embodiment, it is possible toobtain the same effect as that in the first embodiment even in aradiography apparatus including one radiation detector.

In each of the above-described embodiments, scattered ray componentsincluded in each of two radiographic images captured by a series ofimaging processes are corrected. However, the invention is not limitedthereto. For example, scattered ray components included in oneradiographic image captured by a series of imaging processes may becorrected and the corrected radiographic image may be used as an imagefor diagnosis. In this case, for example, in the first embodiment, theradiographic image generated by the first radiation detector 20A isused.

In each of the above-described embodiments, the case in which logconversion is performed for each of the values of the correspondingpixels of the first radiographic image data and the second radiographicimage data and the difference between the pixel values is calculated toderive the ratio of the values of the pixels has been described.However, the invention is not limited thereto. For example, each of thevalues of the corresponding pixels of the first radiographic image dataand the second radiographic image data may be multiplied by a weightingcoefficient, log conversion may be performed for each pixel value, andthe difference between the pixel values may be calculated to derive theratio of the values of the pixels. In this case, the weightingcoefficient may be a value that is obtained in advance as a coefficientfor accurately deriving bone density by, for example, experiments usingthe actual radiography apparatus 16. For example, in a case in which theimaging part includes a region (for example, a region corresponding tothe intestinal canal) including gas, such as the abdomen, a weightingcoefficient for removing the pixel value of the region including gas maybe used.

In each of the above-described embodiments, the bone density derivationprocess performed by the console 18 may be performed by the control unit58A or the control unit 58B of the radiography apparatus 16. Inaddition, in a case in which the radiography apparatus 16 includes anoverall control unit that controls the overall operation of the controlunit 58A and the control unit 58B, the overall control unit may performthe bone density derivation process. Furthermore, for example, aninformation processing apparatus that is connected to the console 18through the network may perform the bone density derivation process.

In the first embodiment, the case in which an indirect-conversion-typeradiation detector that converts radiation into light and converts theconverted light into charge is applied to both the first radiationdetector 20A and the second radiation detector 20B has been described.However, the invention is not limited thereto. For example, adirect-conversion-type radiation detector that directly convertsradiation into charge may be applied to at least one of the firstradiation detector 20A or the second radiation detector 20B. Inaddition, for example, a conversion layer that absorbs radiation andconverts the radiation into charge in the direct-conversion-typeradiation detector is made of amorphous selenium (a-Se) and crystallinecadmium telluride (CdTe).

In the first embodiment, the case in which the ISS radiation detectorsin which the radiation R is incident from the TFT substrates 30A and 30Bare applied to the first radiation detector 20A and the second radiationdetector 20B, respectively, has been described. However, the inventionis not limited thereto. For example, a penetration side sampling (PSS)radiation detector in which the radiation R is incident from thescintillator 22A or 22B may be applied to at least one of the firstradiation detector 20A or the second radiation detector 20B.

In each of the above-described embodiments, the case in which bonedensity is derived using the first radiographic image data and thesecond radiographic image data has been described. However, theinvention is not limited thereto. For example, bone mineral content orboth bone density and bone mineral content may be derived using thefirst radiographic image data and the second radiographic image data.

In each of the above-described embodiments, various processes performedby the execution of software (program) by the CPU may be performed byvarious processors other than the CPU. In this case, examples of theprocessor include a programmable logic device (PLD) whose circuitconfiguration can be changed after manufacture, such as afield-programmable gate array (FPGA), and a dedicated electric circuit,such as an application specific integrated circuit (ASIC), which is aprocessor having a dedicated circuit configuration designed to perform aspecific process. In addition, the various processes may be performed byone of the various processors or may be performed by a combination oftwo or more processors of the same type or different types (for example,a combination of a plurality of FPGAs and a combination of a CPU and anFPGA). Specifically, the hardware structure of the various processors isan electric circuit obtained by combining circuit elements such assemiconductor elements.

In each of the above-described embodiments, the aspect in which theoverall imaging processing program is stored (installed) in the storageunit 86 in advance has been described. However, the invention is notlimited thereto. The overall imaging processing program may be recordedon a recording medium, such as a compact disk read only memory (CD-ROM),a digital versatile disk read only memory (DVD-ROM), or a universalserial bus (USB) memory, and then provided. In addition, the overallimaging processing program may be downloaded from an external apparatusthrough the network.

In each of the above-described embodiments, the aspect in which theindividual imaging processing program is stored in the ROM of the memory62 in the control unit 58A (control unit 58B) in advance has beendescribed. However, the invention is not limited thereto. The individualimaging processing program may be recorded on the recording medium andthen provided. In addition, the individual imaging processing programmay be downloaded from an external apparatus through the network.

What is claimed is:
 1. An image processing apparatus comprising: anacquisition unit that acquires a radiographic image generated by aradiation detector irradiated with radiation from a radiographyapparatus including the radiation detector in which a plurality ofpixels, each of which includes a conversion element that generates alarger amount of charge as it is irradiated with a larger amount ofradiation, are arranged; and a correction unit that corrects scatteredray components caused by scattered rays of the radiation included in theradiographic image, using correction data for correcting scattered rayswhich is associated with each of a plurality of regions in theradiographic image.
 2. The image processing apparatus according to claim1, wherein the radiography apparatus includes two radiation detectors,that is, a first radiation detector and a second radiation detector, thefirst radiation detector and the second radiation detector are arrangedalong a direction in which the radiation is emitted, the acquisitionunit acquires a first radiographic image generated by the firstradiation detector irradiated with radiation with a first energy leveland a second radiographic image generated by the second radiationdetector irradiated with radiation with a second energy level differentfrom the first energy level, and the correction unit corrects thescattered ray components included in the first radiographic image, usingeach of the plurality of regions and first correction data associatedwith the first radiation detector, and corrects the scattered raycomponents included in the second radiographic image, using each of theplurality of regions and second correction data which is associated withthe second radiation detector and is different from the first correctiondata.
 3. The image processing apparatus according to claim 2, furthercomprising: a derivation unit that derives at least one of bone densityor bone mineral content, using the first radiographic image and thesecond radiographic image corrected by the correction unit.
 4. The imageprocessing apparatus according to claim 1, wherein the acquisition unitacquires a first radiographic image generated by the radiation detectorirradiated with radiation with a first energy level and a secondradiographic image generated by the radiation detector irradiated withradiation with a second energy level different from the first energylevel, and the correction unit corrects the scattered ray componentsincluded in the first radiographic image, using each of the plurality ofregions and first correction data associated with the first energylevel, and corrects the scattered ray components included in the secondradiographic image, using each of the plurality of regions and secondcorrection data which is associated with the second energy level and isdifferent from the first correction data.
 5. The image processingapparatus according to claim 4, further comprising: a derivation unitthat derives at least one of bone density or bone mineral content, usingthe first radiographic image and the second radiographic image correctedby the correction unit.
 6. The image processing apparatus according toclaim 1, wherein the correction data includes information indicatingintensity of the scattered rays and information indicating spread of thescattered rays.
 7. The image processing apparatus according to claim 1,wherein the plurality of regions include a directly irradiated regionthat is directly irradiated with the radiation and a subject region thatis irradiated with the radiation through the subject in the radiographicimage.
 8. The image processing apparatus according to claim 7, whereinthe subject region includes regions corresponding to a plurality ofparts of the subject.
 9. The image processing apparatus according toclaim 7, wherein the correction data associated with the subject regionis further associated with a body thickness of the subject.
 10. Theimage processing apparatus according to claim 9, further comprising: anestimation unit that estimates the body thickness of the subject from apixel value of the subject region in the radiographic image.
 11. Theimage processing apparatus according to claim 1, wherein the correctiondata is further associated with imaging conditions.
 12. The imageprocessing apparatus according to claim 11, wherein the imagingconditions include at least one of a material forming a bulb of aradiation source, a tube voltage, a material forming a radiationlimitation member, characteristics of a grid, a distance from theradiation source to a radiation detection surface of the radiographyapparatus, or a quality of a material forming a radiation incidentsurface of a case accommodating the radiography apparatus.
 13. The imageprocessing apparatus according to claim 1, wherein the correction unitfurther corrects the scattered ray components included in theradiographic image, using the correction data associated with a distanceof a gap between a surface of the subject which faces the radiographyapparatus and the radiation incident surface of the case accommodatingthe radiography apparatus.
 14. The image processing apparatus accordingto claim 13, further comprising: a derivation unit that derives a firstbody thickness of the subject in a portion of the subject which comesinto contact with the radiation incident surface of the case and asecond body thickness of the subject in a portion in which the gap ispresent, using a pixel value of a region corresponding to the subject inthe radiographic image, and derives the distance of the gap, using thefirst body thickness and the second body thickness, assuming that thebody thickness of the subject is symmetrically reduced.
 15. The imageprocessing apparatus according to claim 2, wherein each of the first andsecond radiation detectors comprises a light emitting layer that isirradiated with the radiation and emits light, the plurality of pixelsof each of the first and second radiation detectors receive the light,generate the charge, and accumulate the charge, and the light emittinglayer of one of the first and second radiation detectors which isprovided on an incident side of the radiation includes CsI and the lightemitting layer of the other radiation detector includes GOS.
 16. Aradiography system comprising: the image processing apparatus accordingto claim 1; and a radiography apparatus that outputs a radiographicimage to the image processing apparatus.
 17. An image processing methodcomprising: acquiring a radiographic image generated by a radiationdetector irradiated with radiation from a radiography apparatusincluding the radiation detector in which a plurality of pixels, each ofwhich includes a conversion element that generates a larger amount ofcharge as it is irradiated with a larger amount of radiation, arearranged; and correcting scattered ray components caused by scatteredrays of the radiation included in the radiographic image, usingcorrection data for correcting scattered rays which is associated witheach of a plurality of regions in the radiographic image.
 18. Anon-transitory storage medium storing a program that causes a computerto perform an image processing, the image processing comprising:acquiring a radiographic image generated by a radiation detectorirradiated with radiation from a radiography apparatus including theradiation detector in which a plurality of pixels, each of whichincludes a conversion element that generates a larger amount of chargeas it is irradiated with a larger amount of radiation, are arranged; andcorrecting scattered ray components caused by scattered rays of theradiation included in the radiographic image, using correction data forcorrecting scattered rays which is associated with each of a pluralityof regions in the radiographic image.